Embodiments of the invention relate generally to positron emission tomography (PET) and magnetic resonance (MR) imaging, and more specifically, to a hybrid PET-MR system having an RF coil assembly and power coil arrangement for providing power to the RF coil that minimizes loss and attenuation of PET signals.
PET imaging involves the creation of tomographic images of positron emitting radionuclides in a subject of interest. A radionuclide-labeled agent is administered to a subject positioned within a detector ring. As the radionuclides decay, positively charged particles known as “positrons” are emitted therefrom. As these positrons travel through the tissues of the subject, they lose kinetic energy and ultimately collide with an electron, resulting in mutual annihilation. The positron annihilation results in a pair of oppositely-directed gamma rays being emitted at approximately 511 keV.
It is these gamma rays that are detected by the scintillators of the detector ring. When struck by a gamma ray, each scintillator illuminates, activating a photovoltaic component, such as a photodiode. The signals from the photovoltaics are processed as incidences of gamma rays. When two gamma rays strike oppositely positioned scintillators at approximately the same time, a coincidence is registered. Data sorting units process the coincidences to determine which are true coincidence events and sort out data representing deadtimes and single gamma ray detections. The coincidence events are binned and integrated to form frames of PET data which may be reconstructed into images depicting the distribution of the radionuclide-labeled agent and/or metabolites thereof in the subject.
MR imaging involves the use of magnetic fields and excitation pulses to detect the free induction decay of nuclei having net spins. When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but process about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a RF magnetic field (excitation field B1) which is in the x-y plane, i.e. perpendicular to the DC magnetic field (B0) direction, and which is near the Larmor frequency, the net aligned moment, or “longitudinal magnetization”, MZ, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt. A signal is emitted by the excited spins after the excitation signal B1 is terminated and this signal may be received and processed to form an image.
When utilizing these signals to produce images, magnetic field gradients (Gx, Gy, and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
In combination PET-MR systems, the RF coil (i.e., whole body RF coil) that generates the RF magnetic field is typically driven by power cables, also referred to as drive cables. These RF power cables may be as thick as 8 mm in diameter or greater and carry ˜30-35 kW of RF power. The RF power cables are typically mounted at the central section of the RF body coil, which is the virtual ground of the RF coil, in order to minimize the shield currents present on the RF power cables. However, mounting the RF power cables at the central section of the RF body coil leads to significant loss/attenuation of PET signals (measured at ˜15%), which in turn affects PET image quality.
Other techniques to minimize the shield currents on the cables, such as implementing a quarter wave sleeve and/or employing stub baluns, may be used. However, such workarounds to the issue of shield currents on the RF power cables detrimentally requires much wider bore space, which can pose other significant challenges such as the redesign of the magnet/gradient coil.
It would therefore be desirable to design an RF power cable arrangement that provides power to the RF coil but that minimizes the loss/attenuation of PET singles in the PET-MR system. It would also be desirable to minimize the shield currents present on the RF power cable without the need for additional shielding and/or redesign of the magnet/gradient coil.